Definitions
As used herein, “Heliox” is defined as a gas mixture of helium and oxygen whose physical properties are summarized in Table 1 depending on the concentration of Helium.
TABLE 1Physical properties of Heliox at 273 K,1 atmosphere.Percentage of Helium020406080100Density (g/L)1.4291.1790.9290.6790.4290.179Viscosity (μP)204201.2198.4195.6192.8190Kinematic14.317.121.428.844.9106.1Viscosity(μm2 · s−1)
As used herein, “ambient air” is defined as that air which normally exists around us which is either inhaled and exhaled from the environment, or, pumped into a mechanical hand held device from the environment and then inhaled.
As used herein, “aerosolization” is primarily defined as the generation and then breakup of a liquid sheet into primary and satellite droplets, generally 1 micron to 20 microns in size, although the physical form of particles in an aerosol as used herein may be liquid drops or solid dry powder particles.
As used herein, “fluidization” is defined as the deagglomeration of a compact mass of drug in micronized dry powder form manufactured with a preferred particle size range of 1 micron to 5 microns into a cloud, with the objective being the generation of particles in the preferred 1-10 micron range, and more preferably in the 1-3 micron range.
As used herein, “heterodisperse aerosol” or “heterodisperse particle cloud” shall be defined as a deliverable form of a liquid drug formulation or dry powder drug formulation, such that there are particles of many different sizes.
As used herein, “monodisperse aerosol” or “monodisperse particle cloud” shall be defined as a deliverable form of a liquid drug formulation or dry powder drug formulation, such that the particles are all the same, or very near the same size.
As used herein, “alveoli” are air sacs deep in the lung at the terminal end of the smallest and last branch of bronchioles, where gas exchange takes place between the airspace in the lungs and arterial blood. Small particulate drug matter can enter the alveolar spaces, depending on their size and deposition characteristics. After entering the alveoli, the drug matter becomes engulfed by alveolar macrophages, which exist around each alveolus under its surfactant layer and enter the acinus by way of the terminal bronchiolar lumen. Drug particles may be absorbed from the lung primarily by alveolar macrophages.
As used herein, “fine particle dose” shall mean particles that are preferably about 5 μm or less, generally 3 μm or less, and more preferably 2 μm or less.
As used herein, “respirable fraction” (RF) is a dose fraction of aerosolized drug particles small enough in diameter to escape the filtration machinery of the airways and be deposited in the lungs.
As used herein, the terms “dry powder formulation” and “liquid formulation” are pharmacologically active drug by itself, or with any of the following including but not limited to propellants, carriers, excipients, surfactants, anti-microbial, flavoring, and other additions to the formulation that enhance production, shelf life stability, generation of particles, delivery to the desired site in the lungs, and absorption, macrophage or other processed base transfer from the air space into the tissue and blood, or taste.
General Medical Background
Delivery of therapeutic drugs via the lungs for respiratory and non-respiratory systemic diseases, is increasingly being recognized as a viable if not superior alternative to administration of drugs orally/nasally, rectally, transdermally, by intravenous needle injection, intra-muscular needle injection, or gas jet driven non-needle injection through the skin and into the muscle.
Around 1 million patients in the US receive intravenous morphine for the relief of chronic and terminal pain. Morphine actually acts more rapidly with respect to pain management when inhaled than when injected. In addition, there is a major effort to move away from CFC or other vapor pressure based propellant driven inhalers toward alternative technology, due to environmental issues.
All but oral and rectal modes of administration, ideally require a liquid form of drug. Hard particulate drug forms are being explored for gas jet driven needle-less injection through the skin for deposit into the muscle for extended or timed release of the drug substance.
In each of these non-pulmonary methods of drug administration, far higher doses of drug substance than that required for actual therapeutic effectiveness on the target system must be administered to assure that the required therapeutic amount of drug substance is actually delivered to the target system or site. This represents a risk factor to the patient, in that there is a therapeutic variable regarding the amount of dose delivered to the target system or site. The exception is where that target is very local to the site of administration (i.e., mouth, colon, patch of skin, area of muscle, etc).
In addition, many new drugs being developed by companies in the biotechnology field based on peptides and proteins, exist as dry powder in their optimum and/or most stable form, and so these drugs cannot be injected using a needle or needle-less method, or administered transdermally. Genetically produced peptide and protein based drugs are also very sensitive to being altered by in-vivo environmental factors such as enzymes and acids.
If such sensitive drug molecules in dry powder form are delivered orally, they are subjected to the enzymes and acids in the digestive tract. This can reduce the quantity of these sensitive therapeutic molecules available for absorption into the blood in their original therapeutic structure, increasing the need to initially deliver a higher oral dose. Rectal drug administration is neither pleasant, socially acceptable, or commercially viable except in extreme cases where no other choice exists.
The intravenous needle method of administering therapeutic drugs in liquid form in the arm or femorally, results in the dilution and loss of administered drug potency as the blood passes through the venous system back to the heart, then to the lungs, and finally into the arterial circulation for delivery.
Intra-muscular needle injection adds a pathway where part of the administered dose can be lost. The same is true for a gas driven jet needle-less injection, where the drug substance must go through the skin, into the muscle, (usually and primarily) into the venous blood system, and then into the arterial system.
Hence, it is necessary to inject more drugs, regardless of the method, than is really needed to achieve the desired therapeutic effect on, for example, a specific organ system or organ based receptor target fed by arterial blood. However, by introducing a drug substance into the arterial blood stream at its source, the lungs, a bolus of drug delivered to the target is less diluted, and, therefore, less drug needs to be deposited in-vivo at the site or entry point of administration (the alveoli).
Delivery of drugs via the lungs is the optimal approach to treat diseases in the lung. In addition, drugs delivered via the lungs for other than respiratory diseases, go rapidly and directly into the arterial blood, then to the heart, and then to the other critical organs such as brain, liver and kidneys, and receptor sites residing thereon. This reduces the effect of dilution on the administered therapeutic dose in the bloodstream. Furthermore, there is minimal enzymatic or acid activity in the lungs compared to the stomach that can impact the therapeutic molecular integrity of sensitive drug molecules such as genetically engineered peptides and proteins. Pulmonary drug delivery can, depending on the drug and disease:                a) improve the efficacy of a drug;        b) improve the bioavailability of a drug, which is particularly important for biological compounds such as peptides and proteins;        c) improve targeting to an organ or receptor site thus reducing unwanted side effects (which is an important consideration with, for example, anticancer agents); and        d) mimic the biopattern of a disease, or circadian rhythm, e.g., as in the case of sustained-release anti-hypertensives designed to peak coinciding with the early morning blood pressure surge.        
Commercially, a new method of pulmonary drug delivery for an existing drug, can extend its therapeutic indications, lower cost, and facilitate a more rapid time to market. Since drugs administered by the pulmonary route do not require sterility, a sterile device or sterile environment, they are ideal for the delivery of drugs in difficult environments.
U.S. Pat. No. 6,125,844 discloses an apparatus for portable gas-assisted dispensing of medication not using a fluorocarbon propellant. The apparatus comprises a pressurized supply of gas containing therapeutic gas or mixture of therapeutic gases, and one or more drugs mixed therein, connected to a pressure regulator, wherein the pressure regulator is connected to a gas release switch which is connected to a breath activator. The breath activator is connected to an aspiration chamber, whereby in use when a patient inhales from the aspiration chamber, the inhalation causes the breath activator to engage with the gas release switch to release the therapeutic gas/drug mixture into the aspiration chamber, wherein the therapeutic gas and drug in the aspiration chamber are simultaneously delivered to a patient during inhalation. Alternatively, medication can be stored in a separate drug reservoir adjacent the pressurized supply of therapeutic gas, which medication is drawn into the aspiration chamber by a venturi assembly.
Variables that affect inhaler generated particulate drugs being delivered to the right location routinely mentioned in the medical literature include:
a) those that are breathing related including the volume of inspiration, inspiration flow rate (velocity), breath holding period after inspiration of a dose, the total lung volume at the time the bolus of medication is administered, and the expiration flow rate;
b) those that are particulate related including aerosol particle size, shape, density of the liquid or powder drug particles, and size distribution in the dry powder or liquid aerosol cloud produced; and
c) the medical status of the patient, and in particular, the status of the respiratory system of the patient.
The objective with any method and technology involving inhalers, is: a) to generate particles of the optimum size range for deep lung delivery, and b) to get any administered particles past the larger airways where they will be lost to turbulence and impaction and into the middle (for treating respiratory diseases) and deep (for delivering drugs to the target area where they can enter the arterial blood) lung.
Unlike intravenously administered drugs, drugs administered via the lungs are not subject to prior first pass hepatic metabolism. They are also less subject to reacting with or being affected by fewer receptors prior to reaching their intended target either in the lungs or systemically, resulting in a reduced amount of drug being needed, if the particle size and delivery to the target location in the lungs are optimized. However, because any systemic drug administered by the lung does go straight to the heart first, the cardiac side effects of excipients and drugs administered by this method are an issue. As an example of the rapid effects drugs administered via the lungs can have systemically, administration of the pain killer morphine via the lungs is faster acting than morphine administered intravenously.
Recognition of the ability to deliver systemic therapeutic drugs by inhalation due to the physiological properties of the lung and circulatory system, has led to a large number of different therapeutic drugs being developed and evaluated for administration by inhalation to treat even non-respiratory diseases.
A key problem is in the maximizing the number of these smallest particles that are delivered to the terminal branches of the bronchioles and the alveoli. Small particles, preferably 1 μm-3 μm in size, are optimal for this purpose. Generally, only about 10-20% of the amount of particulate drug dispensed by conventional inhalers is delivered in this range.
Large molecule drugs, such as peptides and proteins which are now possible due to genetic engineering, do not pass easily through the airway surface because it is lined with a ciliated mucus-covered cell layer of some depth, making it highly impermeable. The alveoli however, have a thin single cellular layer enabling absorption into the bloodstream. The alveoli are the door to the arterial blood and are at the base of the lungs.
So, to reach the alveoli, a particulate drug must be administered in small size particles, and the inhalation must be moderated, slow, and deep. Large particles will impact in the oropharyngeal area or settle in the upper bronchi. If the particles are too small and/or ultra light, they will be exhaled (the latter is especially true if air is the tidal front of gas entraining the ultra light particles).
The larger passages through which the air and drug particles travel generates turbulence, which also results in the impaction and loss of drug particles. A desired goal is to increase the laminar flow of the gas stream in the larger air passages, so that particles reach the smaller passages where laminar flow is naturally induced. If there are any constrictions in the bronchi or bronchioles, resulting, for example, from asthma, the turbulence and rate of impaction of drug particles can also increase at those points of constriction.
Any variability in the dose deposited in the lungs, and where it is deposited in the lungs, could have a major effect on treatment because of the narrow therapeutic range of many drugs, and the potency of such drugs. One well known such example is insulin.
Aerosol particles are deposited in the airways by gravitational sedimentation, inertial impaction, and diffusion. All three mechanisms act simultaneously. However, the first two are the principle methods that apply to the deposition of large particles. Diffusion, is the primary factor of deposition of smaller particles in peripheral regions of the lung.
The optimum size particles of drug for delivery to the alveoli are in the range generally of 1-3 microns, and usually particles less than 2 microns reach the alveoli.
The diameter of therapeutically usable particles is generally between 0.5 and 5 microns. Particles 1-5 microns are deposited in the larger airways while particles generally below 3 microns in diameter reach the terminal bronchioles and alveoli and are optimal for transference into the arterial blood. The depth of penetration of a particle into the bronchial tree is inversely proportional to the size of the particle, down to 1 μm. Particles smaller than 1 μm, however, are so light that a large proportion does not deposit in the lungs.
The small airways are the optimal sites for the inhalative treatment of obstructive pulmonary diseases. Diffusion is a process that applies to particles smaller than about 3 microns. The maximum collection of particles by the deep lung is by the process of sedimentation.
Some of the sub-micron particles of a drug may be exhaled because their sedimentation may not be high enough in air—which is normally the ambient entrainment gas and environment in the lungs.
Prior art, whether metered dose inhalers (MDI) or dry powder inhalers (DPI), use air as the exclusive or primary means of conveying fluidized powder or aerosolized liquid drug into the lungs. In the case of MDIs, it is assumed that the propellant evaporates as intended or constitutes a very small fraction of the total gas inhaled at full tidal volume with the drug dose and air.
Heliox has been administered to a patient in a hospital setting prior to the administration of a dry powder or liquid aerosol drug. Heliox has also been used to administer a liquid drug using a nebulizer, which is a different type of device for pulmonary drug administration lasting 10-60 minutes. That is distinct from “puffs” received through an inhaler. Additionally, in both cases, the systems in which Heliox were used were designed for the physical properties of air and not Heliox, and so were not optimized for Heliox.
Prior art and medical publications pertaining to inhalers, address other factors but do not focus on the specific gas involved in the transport of particles into the lung. In the case of DPIs, the gas is always assumed to be, or stated specifically to be, air. In the case of MDIs, the “gas” is always assumed to be a liquid propellant having a vapor pressure, CFC in most cases, and is only a negligible fraction of the inhaled volume, the balance being air.
MDI is a metered dose inhaler consisting of a propellant generating a vapor pressure and a drug in suspension or solution form, where, when the device is activated, the vapor pressure of said propellant pushes a predetermined amount of liquid drug through a nozzle generating an aerosol for inhalation. MDIs contain suspensions or solutions of a drug, a propellant, and a surfactant that acts as a lubricant to stop particles from aggregating and to reduce clogging of the aerosol nozzle. MDIs rely on the use of propellants that have a high vapor pressure. The higher the vapor pressure, the faster a liquid containing a drug can be pushed out of a nozzle, and thus a thinner liquid sheet is formed, and smaller particles are produced. Vapor pressure is therefore directly related to the velocity generated and the fraction of fine or desirable small particles generated.
Pressurized aerosols historically used chlorofluorocarbon (or CFC) propellants generating a pressure of approximately 400 kPa or higher. The aerosol cloud therefore emerges from the canister at a high speed. Furthermore, the drug crystals are initially enclosed within large propellant droplets whose mass median diameter may exceed 30 μm. Large particles traveling at high velocities are very susceptible to oropharyngeal deposition by inertial impaction. While the propellant evaporates and the particles slow down when the device is held away from the mouth, or when an MDI spacer is used, on average, only about 20% of the original or nominal dose actually enters the lungs.
In an MDI, the generation of an aerosol occurs in what can only be described as an explosive manner since the propellant containing the therapeutic solution or suspension disintegrates as it passes through the aerosol nozzle at very high velocity. As the propellant flash rapidly evaporates, the liquid particles decrease rapidly in diameter to the state of a “dried solute”.
The velocity of the discharged particles entrains the evaporating particles as they exit the device and move into the airstream. This velocity is much higher than an inhalation velocity by a user. The result can be impaction of particles in the oropharyngeal area. A spacer, which is discussed later, is a solution to this problem, i.e., reducing the velocity of the “cloud” of particles prior to inhalation. Another technique is to use the “open-mouth” method that implies activating the device a few cms away from an open mouth.
MDIs containing a suspension require that they be shaken before use. MDIs containing a solution need not be. This presents a problem to patients using more than one type of drug, i.e., one in suspension and one in solution, as the patient may shake the wrong MDI, or not shake the MDI that needs to be shaken before use. The latter one would result in an incorrect dose of the drug being delivered and inhaled. This is an advantage to the use of DPIs, as there is no “to shake or not to shake” decision. MDIs containing propellant and a suspension or solution, also present a challenge concerning stability over a temperature range.
A problem with both MDIs and DPIs is that there is often poor coordination between the patient pressing the actuator and the timing of the inhalation. One solution is to use a spacer between the device and patient, that will also allow for the heavier particles to settle before the patient inhales.
Another problem with MDIs is that they are based on propellants that rely on vaporization to generate pressure, and a drop in temperature occurs when vaporization occurs. The vaporized propellant can hit the back of a user's throat before it has completely evaporated if no spacer is used. This can lead to reflex gagging which interrupts the continuous and deep inhalation required for optimum delivery of the drug. In addition, water moisture in the mouth will condense rapidly in the cold vapor, causing the small liquid medication droplets to coagulate and drop out, reducing the percentage of drug actually deliverable past the oropharyngeal area.
DPI is a dry powder inhaler consisting of a drug in micronized dry powder form provided in a compact shape and contained in a unit dose container or reservoir, which is fluidized by the flow of a gas and inhaled by the patient.
Micronized dry powder formulations are very soluble and quickly dissolve in the fluid layer on the surface of the deep lung before passing through the thin single cellular layer of the alveoli. They are then deposited in the alveolar region and can be absorbed into the bloodstream without using what are commonly referred to as penetration enhancers. Dry powder aerosols can carry approximately five times more drug in a single breath than metered dose inhaler (MDI) systems and many more times than liquid or nebulizer systems.
Micronized dry powder drugs used in inhalers are usually produced with an original particle range of 1-10 microns. An individual dose as loaded can take from 5 mg to 20 mg of dry powder drug. A lower total amount of dry powdered drug is possible with purer drugs, or with drugs that do not require or are packageable without excipients. Examples of excipient carriers used in dry powder drug formulations include lactose, trehalose, or crystalline or non-crystalline mannitol. Trehalose and mannitol, which are spray dried sugars, are better dispersal agents than lactose.
Thus, the “drug substance” in a DPI consists of the pure drug, plus a sugar if an excipient is used, compared to the multitude of constituents contained in a MDI. This multitude of constituents in a MDI increases the work involved in production of the product and its packaging, can effect formulation stability, can cause aerosolization problems by clogging the nozzle, and may require either the shaking or non-shaking of the MDI Inhaler before use.
In DPI devices, providing compressed gas or propeller/impeller assisted fluidization, basing the fluidization on the patient's inhalation produces a major variability in dosing and particle size formation. The velocity, ramp up rate, and continuous event of this inhalation are variables that can effect the fluidization of the powdered drug and the effective delivery of the optimum size particles to the deep lung. The higher the rate of gas velocity, the finer the particle size created during fluidization, but the greater the possibility for impaction of particles in the oropharyngeal area during inhalation, where the gas velocity which fluidizes the dry powder drug is derived from the “suction” or negative pressure of a strong inhalation.
Devices that rely on the force of the patients inhalation, also operate based on the “suction” or pulling effect of said gas flow, i.e. a negative pressure, to pull apart and fluidize the drug powder. This is less effective than a highly focused directed stream of high pressure gas, which is consistently delivered at the same pressure.
Some DPIs use compressed air generated by a pumping mechanism, which the patient utilizes, whereby the pressure is released for fluidization of the powder drug when the system is actuated. The pressure, and therefore velocity, of a gas that can be generated by a hand pump or an inhaler device, is far less than that available from a compressed gas cartridge. The uniformity of fluidization of the dry powder would therefore be less using a manual hand pump, with the possibility therefore of generating larger percentages of larger size particles, which result in the variable and inconsistent loss of drug in the oropharyngeal and upper bronchi.
The higher the velocity of the gas hitting the dry powder, the greater the amount of powder dislodged and the turbulence induced, which can create a cloud of particles for inhalation. In the case of dry powder inhalers, the ramp speed to the velocity required to deaggregate or deagglomerate the dry powder into fine particles, is as important a factor as velocity in determining effectiveness.
Systems using dry powder drug in capsules, require the patient to load the capsules individually, whether the system is capable of being loaded with one dose at a time, or several doses for multi dose use over time. In some of these devices, the capsule is crushed to thereby release the powder contained therein.
A DPI entrains the fluidized drug powder and sends it through a narrow gap, increasing the velocity of the gas and powder to improve deagglomeration by turbulence and reduce the number of large particles by impaction or settling out. Often, a baffle is also included in the system to trap larger particles.
One problem in using compressed gas vs. a hand pump to generate compressed air DPI (or a liquid MDI driven by CFC vapor pressure) is that the compressed gas pressure will decrease with usage. In the case of the hand pump driven DPI, the gas pressure is consistent during each dose fluidization procedure. In the case of gas driven MDIs, the pressure available for aerosolization decreases over time near the end of the capacity, unless the MDI has a cut off which does not allow dose administration below a certain minimal pressure required to achieve sufficient aerosolization.
A spacer is a plastic or metal tubelike device that is placed between the inhaler device and the patient, and into which the inhaler device delivers the particulate cloud generated by dry powder fluidization or liquid aerosolization. A spacer can be open-ended, allowing a slowing down of the gas, or closed-ended (holding chamber) to reduce the loss of dose inhaled due to poor hand-breath coordination. The spacer slows down the gas mass and particles leaving the inhaler, traps larger particles by impaction and settling, and provides a better control of inhalation rate and timing, delivery of the desired size range of particles, and reduced oropharyngeal loss of particles due to impaction, versus inhaling directly from the inhaler device. It also reduces the gagging effect from inhaling a cold gas like Freon. Spacers have been incorporated into the routine use of MDIs.
Inhalation flow velocity in inhalation driven inhaler determines the quality of the aerosol cloud, as the greater velocity fluidizing the dry powder drug, the finer the particles produced. However, the inhalation of particles at a fast rate, leads to impaction of a large percentage of particles on the back of the throat.
Heliox, which is commercially available in a combination of 70% or 80% helium in oxygen, has been used for over 70 years in respiratory therapy. Heliox is administered in some hospitals and emergency rooms in large gas cylinders. The most popular types are the “K” cylinder that stands 51 inches in height, 9 inches in diameter and weighs 130 lb when fully filled. Heliox is supplied at 2,200 psig and requires a two-stage pressure regulator to reduce the pressure for administering to patients. However, due to its bulkiness and requirement of sophisticated pressure and flow regulators, it is used only in research and hospital facilities.
Gas flow within the tracheobronchial tree is complex and depends on many factors. For a given pressure gradient, the volumetric flow rate of a gas is inversely proportional to the square root of its density. In accordance with the subject invention, it has been found that substituting helium for nitrogen in inhaled gas mixtures results in increased gas flow rates because the density of helium is much lower than that of nitrogen.
Resistance to the flow of gas within the tracheobronchial tree results from convective acceleration and friction. Convective acceleration is the increase in the linear velocity of fluid molecules in a system of flow in which the cross-sectional area is decreasing. Frictional resistance may be either turbulent or laminar depending on the nature of the flow. Since resistance associated with these factors is density-dependent, breathing a less dense gas should decrease flow resistance and, consequently, reduce respiratory work. An obstruction in the upper airway causes a resistance to flow that is primarily convective and turbulent and therefore susceptible to modulation through a change in gas density. For respiratory treatment, it is desirable to create a flow of minimum pressure drop or flow-resistance.
Gas flow in airways may be laminar, turbulent, or a combination of the two. Turbulence is predicted by a high Reynolds number, which is a unitless quantity proportional to the product of gas velocity, airway diameter, and gas density divided by viscosity. The Reynolds number is also expressed as the ratio of kinetic to viscous forces. The decreased density of helium, when substituted for nitrogen, lowers the Reynolds number and may convert turbulent flow to laminar in various parts of the airway. Turbulence is highly dependent on the surface roughness, so that a flow in a rough cavity might be turbulent even if the Reynolds number predicts a laminar flow. Even in the absence of turbulent flow, the decreased density of helium improves flow and decreases work of breathing along broncho-constricted airways.
The efficacy of Heliox in respiratory therapy occurs because it is a low-density gas. The rate of diffusion of a gas through a narrow orifice is inversely proportional to the square root of its density (Graham's Law). When an area of stenosis occurs in the airway, there is resistance to flow at the site of the stenosis. The resistance varies directly with gas density. Downstream from the stenosis, airflow becomes turbulent. By substituting helium for nitrogen in inspired air, resistance at stenotic areas is reduced and turbulence downstream from the stenosis is either reduced or eliminated.
In the tracheobronchial tree, a laminar flow normally exists in airways that are generally less than 2 mm in diameter. Turbulent flow has been observed in the upper respiratory tract, the glottis, and the central airways. This upper portion of the airway, especially the throat, and the main bronchioles, are considered to be the region where the turbulent intensity is sensitive to the gas density.
Since airway resistance in turbulent flow is directly related to the density of the gas, Heliox, with its lower density than nitrogen or oxygen, results in lower airway resistance. Heliox further lowers airway resistance by reducing the Reynolds number, such that some areas of turbulent flow are converted to laminar flow. The higher flow rate of Heliox has the ability to stay laminar at velocities under which air would be turbulent.
Heliox does not need to be laminar to provide higher flow rates and its benefits persist under turbulent conditions. Some have the misconception that, due to its lower density, helium is less viscous than air, so it flows faster. Actually, the absolute viscosity of helium is slightly higher than that of air, and its kinematic viscosity (absolute viscosity divided by density) is about seven times that of air. Thus, from the fluid-dynamical standpoint, helium is more viscous than air.
The linear relationship between helium concentration and resistance to flow is predictable on the basis of fluid mechanics. Helium has two major effects in reducing resistance in an obstructed airway. First, helium reduces the probability of turbulence. Flow of air in the upper airway is turbulent, except at rest, because of the rough walls of the airway and the relatively short lengths of the airway segments compared to their diameters.
The probability of turbulent flow is predicted by the Reynolds number:
                    Re        =                              ρ            ⁢                                                  ⁢            VD                    μ                                    (        1        )            Where                D=Diameter of the mouth, airway or throat (cm)        V=Gas velocity (cm/sec)        ρ=Density of the gas (g/cc)=        μ=Viscosity (g/cm/sec)        
Second, gas flow through an orifice requires an increase in pressure to maintain the flow:
                              U          o                =                                            C              o                                                      1                -                                  β                  4                                                              ⁢                                                    2                ⁢                                  (                                                            P                      a                                        -                                          P                      b                                                        )                                            ρ                                                          (        2        )            
where Pa−Pb is the pressure difference caused by the orifice (dynes/cm2), and Co, is the discharge coefficient, which depends on the sharpness of the edge of the orifice.                U0=Velocity through the orifice        β=Ratio of orifice diameter to pipe diameter        Pa=Pressure at upstream before orifice        Pb=Pressure at downstream after orifice.        ρ=Density        
In summary, Heliox is more beneficial because of its lower density. Compared to air, it flows at a higher flow rate for fixed pressure gradient, or needs a lower pressure gradient or work of breathing (or patient inhalation effort) for a given flow rate. This is valid even in turbulent conditions.
There is medical literature where Heliox has been provided to a patient prior to dosing with an Inhaler based on a CFC based propellant. There is also a study where a small volume of Heliox (40-70 ml) was delivered as bolus but with a shallow breath during pulmonary administration of a particulate to see if the entrained particles would diffuse deeper into the lungs by themselves within the Heliox gas.
There is also literature where Heliox was used with a nebulizer to deliver a drug in liquid form. Most of the time, the velocity of the nebulizer gas flow was based on that used for air. In other cases where the gas flow velocity was altered, the aerosol nozzle used was designed for air and not Heliox or pure helium, so that the particle size distribution was not adapted to the change of gas.
Two factors that can influence the delivery of an optimally fluidized dry powder drug formulation are static electricity and humidity. It is desirable to avoid imparting a static electricity charge to the fine particles, especially those 1 micron or less in size. The static charge will form an attractive force on the particles, causing them to clump together, rendering them of a collective size that is unsuitable for deep lung delivery. This type of particle cohesion is highly undesirable because a few particles that are attracted together can double or triple the terminal settling velocity. This is a key reason why conventional inhalers using inhaled air, propeller driven air, or compressed air pumps, have more than 50% of the drug lost in the mouth and throat, before they can enter the lung.
Moisture in the fluidization gas can also result in the clumping of particles. This is a disadvantage of using inhaled air, air from the surrounding environment driven through a propeller, or air compressed using a hand pump that is part of an inhaler. If an inhaler is used in a humid geographical location or during humid seasonal conditions, the humidity can affect the deliverable dose of drug particles in the size range required for penetration into the deep lung, thereby affecting the dose.
In addition, if moisture comes in contact with the powder before it is fluidized, the moisture can accumulate on the outer layer of the powder, forming lumps before fluidization occurs.
The subject invention system can be light enough to be portable, and small enough for a child up to an adult to hold and use.
It is an object of the subject invention to provide an inhaler that can deliver appropriate sized particles to the lungs efficiently using a propellant with sufficient pressure to fluidize or aerosolize a drug to be used by a patient.